Many thousands of patients are treated every year for heart valve dysfunction. Of these patients, a significant number end up having one or more of their heart valves surgically replaced. In general terms, replacement valves are either xenografts or homografts. Xenografts include mechanical or biological valve prostheses. Homografts include cryo-preserved or glutaraldehyde-fixed explanted valves. Both types of valve have drawbacks.
For example, to avoid thromboembolic complications it is typically necessary for patients fitted with mechanical valve prostheses to be prescribed anticoagulation drugs for life, thereby permanently increasing that patient's risk of haemorrhaging. Infections are a further, often life-threatening complication for the patient. Biological xenografts, for example pig valves treated with glutaraldehyde, can be employed but their tendency to calcify poses a problem.
Homografts, i.e. fixed heart valves isolated from human donors, are relatively resistant to infections, but are nonetheless exogenous tissue which can still cause immune reactions. The availability of homografts tends to be limited, and what's more they also tend to calcify and are therefore subject to considerable degeneration—typically requiring re-operation after 7 to 12 years.
In addition to the disadvantages already described, a further disadvantage is that they are all made of inorganic or fixed organic material and therefore lack the capacity for repair, for reconfiguration or for growth. As a result it is commonplace, in particular for younger patients, to have to undergo a series of operations as they grow up. Aside from the inherent risk associated with any surgical procedure, fusions occurring in the thorax following preceding operations tend to significantly increase the risks associated with each subsequent re-operation.
To alleviate problems such as these it has previously been proposed to “grow” autologous bioengineered prostheses for surgical implantation into patients. Such prostheses, being grown from the patient's own tissue, are less likely to provoke an adverse immune response. Furthermore, as these implants are formed of living tissue, they are able to grow with the patient to such an extent that the need for re-operation (to replace prostheses that have reached the end of their life) is greatly reduced.
Our previous work has proposed an in vitro method whereby an autologous heart valve can be grown in a bioreactor. In this method a biodegradable support is incubated with homologous fibroblasts and/or myofibroblasts to form a connective tissue-like matrix that is then colonized with (typically autologous) endothelial cells. The connective tissue-like matrix is then transferred into a bioreactor for culturing and conditioning of the implant to the flow conditions that it is likely to experience when transplanted into the human body. The scaffold degrades (at least substantially) whilst in the bioreactor, so that at the time of its implantation, the prosthesis is comprised of some biomaterial and preferably autologous cell material that can readily be sewn into the receiving patient's heart. Further details of such prostheses are disclosed in International PCT Publication No. WO 2004/018008, the contents of which are incorporated herein by reference.
FIG. 1 is a schematic representation of a bioreactor system 1 that we have previously proposed for the production and conditioning of heart valves as described above. As depicted, the bioreactor system 1 comprises a bioreactor 3, a reservoir 5 for culture medium, and an eccentric pump 7. An inlet conduit 9 couples an outlet 11 of the reservoir 5 via a non-return valve 13 to an inlet 15 of the bioreactor 3, and an outlet conduit 17 couples an outlet 19 of the bioreactor 3 via a second non-return valve 21 to an inlet 23 of the reservoir 5. As depicted, the reservoir inlet and outlet ports are always submerged in fluid so that the bioreactor 3, inlet and outlet conduits are always substantially free of air. The pump 7 is coupled to the bioreactor 7 via a fluid (typically air) supply line 25, and is operable (as will later be described) to drive fluid into and out of the bioreactor to cause a circulation of culture medium through the system.
The bioreactor is illustrated schematically in more detail in FIG. 2. As shown, the bioreactor 3 comprises an outer casing formed by a bottom section 27, a middle section 29 and a top section 31 bolted together by a series of bolts 33. A resilient membrane 35 is sandwiched between the bottom 27 and middle 29 sections and defines a pumping fluid cavity 37 that is coupled to the pump 7 via inlet/outlet port 39 and fluid line 25 (see FIG. 1). The middle section 29 defines a culture medium inlet chamber 41 that is in fluid communication (by means of a central bore 43) with a nutrient chamber 45 defined by the top section 31.
As depicted, the central bore 43 projects into the nutrient chamber 41 to form a mounting stub 47 onto which mounting ring 49 is mounted. A tissue prosthesis, such as a valve (not shown), can then be fixedly attached to the mounting ring for conditioning in the bioreactor.
In use, the pump 7 is operated to provide a pulsatile flow of fluid through the tissue prosthesis provided in the bioreactor. This pulsatile flow is provided by driving gas (typically air) into and out of the fluid cavity 37 to deform the resilient membrane 35 in an upward (on a pumping stroke) and downward (on a return stroke) direction (as depicted).
During the pumping stroke, gas is pumped into the gas cavity 37 to deform the resilient membrane 35 upwardly. Upward deformation of the resilient membrane drives fluid from the inlet chamber 41 up through the central bore 47, through the tissue prosthesis (not shown) fixedly attached to the mounting ring 49, and out of the top section via outlet port 19. As the membrane 35 moves upwards, the first non-return valve 13 opens to allow fluid to flow from the reservoir outlet port 11 via inlet conduit 9 and into the inlet chamber 41 via inlet port 15. Simultaneously, the second non-return valve 21 opens to allow fluid to flow out of the outlet port 19 and into the reservoir via outlet conduit 17 and reservoir inlet port 23.
The return stroke of the pump 7 draws air out of the gas cavity 37 and deforms the resilient membrane 35 downwardly. Movement of the resilient membrane 35 downwards closes both of the non-return valves and at least substantially prevents fluid flow back through the tissue prosthesis mounted on the mounting ring 49.
As we have previously disclosed, the pump rate should be closely controlled so that the rate of fluid flow through the tissue prosthesis mounted in the bioreactor is slowly increased, principally to avoid damaging the prosthesis, to a point where it mimics that which one might expect to find when the prosthesis is implanted in a patient. This approach provides for a slow conditioning of the prosthesis to the flow rates and pressures that one might expect the prosthesis to experience when implanted in a patient, and as a result renders the prosthesis less likely to failure on eventual implantation into a patient.
Whilst this approach has previously worked well, we have noted that fine control of eccentric pumps is problematic. In particular, we have noted that whilst we can set an eccentric pump to operate at a given speed (and hence theoretically provide a particular output (in terms of liters per minute)), we do not know what the actual pump output is. Furthermore, the nature of the eccentric pump design means that consecutive pumping strokes do not necessarily provide the same output. This problem can be exacerbated at higher pump rates, the effect being that the error in expected output (i.e. the pump output that one would expect for a given pump setting) increases quasi-exponentially as the pumping rate increases. As our prosthesis conditioning process involves operating the pump at indicated rates of up to roughly 5 liters per minute (a relatively high rate) the error between the set pump output and the actual pump output can be considerable, and more importantly liable to fluctuate with each pumping stroke.
A related problem is that whilst it is possible to set the pump to operate at a given indicated rate, we can only infer from this indicated rate the actual rate of culture medium flow through the bioreactor. Given the inaccuracies in pump output, and the fact that we are driving the membrane and not the culture medium, it is likely that the actual rate of culture medium flow is significantly different to that notionally indicated by the setting of the pump. Clearly it would be advantageous for close control of the culture medium flow rate and the conditioning process as a whole for the actual culture medium flow rate to be controllable.
Another problem associated with the bioreactor design we have previously proposed is that it is relatively inconvenient to operate. For example, once a prosthesis has been located in the bioreactor for culturing and conditioning, the bioreactor must be refilled with culture medium. This is done by decoupling the outlet conduit 17 from the second non-return valve 21, and then supplying culture medium (slowly so as to avoid damage to the prosthesis mounted in the bioreactor as might otherwise be caused if the fluid were to be quickly introduced into the chamber) via the inlet 15.
Typically this is done by slowly lifting the reservoir above the bioreactor so that fluid flows under gravity into the bioreactor. The fluid fills the internal voids 41, 45 of the bioreactor until it flows past the second non-return valve 21, and out of the decoupled tubing. At this point the bioreactor is considered to be “filled”, and the reservoir is lowered to cut off the gravity flow of fluid into the reactor. The outlet conduit 17 can then be reconnected to the second non-return valve 21, and the system is ready for operation.
A first problem associated with this arrangement is that as it necessarily involves some spillage of culture medium (through the valve 21 on decoupling of the outlet conduit 17), it is relatively messy—and as such not consistent with desirable clean room procedures. A second problem is that as it is necessary to monitor the amount of culture medium in the bioreactor, a spillage of culture medium means that one does not know exactly how much fluid is left to circulate through the bioreactor. A third problem is one of convenience. As the outlet conduit 17 is full of culture medium one needs one hand to close the conduit while one decouples it from the second non-return valve 21 with the other. Without letting go of the outlet conduit one then needs to lift the reservoir to cause the culture medium to flow into the bioreactor.
A further problem associated with the system depicted in FIG. 1 is related to its size. As shown, the system includes a number of components joined by conduits, and as a consequence the footprint of each system is relatively large. In the context of an industrial scale operation, the relatively large size of each system requires a relatively large amount of space, and this space requirement has ramifications for the cost of the production, and ultimately the sales price of the prostheses produced by the process.